Different methods and measuring devices are known for the tomographic imaging of eye structures, which are based on confocal scanning systems or optical coherence tomography (OCT).
Ophthalmoscopes based on confocal scanners, particularly confocal scanning laser ophthalmoscopes (cSLO), are an important tool for the diagnosis and therapy in ophthalmology (U.S. Pat. No. 6,769,769 B2). Confocal scanners can provide a three-dimensional spatial resolution by limiting the depth of a movable focus using spatial filtering and, contrary to OCT, do not rely on the use of interference effects.
By contrast, in OCT methods, coherent light for distance measurement and imaging is used on reflective and scattering samples using an interferometer. The OCT methods applied to the human eye provide measurable signals during a depth scan due to the changes of the refractive index occurring at the optical boundaries and due to volume scattering. Optical coherence tomography is a very sensitive and quick method for interferometric imaging which has become prevalent in the medical field and fundamental research (Wolfgang Drexler, James G. Fujimoto: “Optical Coherence Tomography Technology and Applications,” Springer Publishing 2008). OCT images (OCT scans) of eye structures are frequently used in ophthalmology for diagnosis and therapy follow-up as well as planning of procedures and the selection of implants. The use of OCT scans of the retina for determining retinal nerve fiber layers (RNFL) for the diagnosis of glaucoma and monitoring of the progress of the disease is one example for the OCT-supported diagnosis.
For example, the basic principle of the OCT method described in U.S. Pat. No. 5,321,501 is based on white light interferometry and compares the travel time of a signal using an interferometer (most commonly a Michelson interferometer). The arm with known optical path length (=reference arm) is used as reference for the measurement arm, which holds the sample. The interference of the signals from both arms yields a pattern for determining the scattering amplitudes on the basis of the optical delays between the arms and thus a scattering profile, which, analogous to ultrasonics, is called A-scan. Quick variations of the optical delay between measurement arm and reference arm can, e.g., be realized using fiber lines (EP 1 337 803 A1) or so-called rapid-scanning optical delays (RSOD) (U.S. Pat. No. 6,654,127 B2). In the multidimensional scanning grid methods, the beam is guided transversely in one or two directions, allowing for the recording of a two-dimensional B-scan or a three-dimensional volume tomogram. When the length of the reference arm is held constant, a two-dimensional C-scan can be obtained with lateral scanning of the measuring beam in two directions.
An important example for the use of optical coherence tomography is the biometry in the eye anterior segment using an anterior chamber OCT (AC-OCT), allowing for the subsequent selection of implants, such as intraocular lens implants (IOL's), particularly also phakic IOL's for a refractive correction. However, IOL's are most commonly used for replacing the natural crystalline lens in case of a clouding of the lens (cataract), wherein refractive and imaging errors are also increasingly corrected.
In addition to OCT, optical coherence domain reflectometry is used for interferometric biometry, which allows for the determination of intraocular distances, which are required as biometric parameters for the selection of IOL's (US 2005/018137 A1, U.S. Pat. No. 7,400,410 B2).
The most important biometric parameters are axis length (distance between cornea and retina), corneal curvature and refractivity, and the depth of the anterior chamber (distance to the crystalline lens). In order to ensure the best possible vision after surgery, it is necessary to determine these biometric parameters with adequately high accuracy. The selection of a suitable replacement lens using the determined measurements is based on established formulas and calculation methods.
The basic principle of the OCT method is based on white light interferometry or short coherence interferometry and interferometrically compares the travel time of a backscattered sample signal (or measuring signal) with a reference signal using an interferometer (most commonly a Michelson interferometer). This allows not only for the analysis of reflections on surfaces but also of slight, varying scattering signals from different sample depths.
The arm with known optical path length (=reference arm) is used as reference for the measurement arm (also called sample arm). The interference of the signals from reference arm and sample arm generates an interference pattern for determining the relative optical path length of scattering signals within an A-scan (depth signal). In one-dimensional scanning grid methods, the beam is then, analogous to ultrasonics, guided transversely in one or two directions, allowing for the recording of a two-dimensional B-scan, a C-scan or a three-dimensional tomogram. Typically, a C-scan is a two-dimensional tomogram, which was obtained through two-dimensional scanning at constant reference arm length in a time-domain OCT. However, in the following, said term shall be used as synonym for all scans which are based on two-dimensional scanning, thus also for volume scans. Here, the amplitude values of the individual A-scans are depicted as linear or logarithmic grey-level or pseudo-color values. It is also known that volume scans can be corrected by comparing them with B-scans with regard to interferences caused by sample movements (U.S. Pat. No. 7,365,856). Furthermore, it is known that a phase-resolved measurement, particularly by Doppler signal analysis, can be used to generate and present additional information about dynamic processes (Adrian H. Bachmann, Martin L. Villiger, Cedric Bluffer, Theo Lasser, and Rainer A. Leitgeb: “Resonant Doppler flow imaging and optical vivisection of retinal blood vessels,” Vol. 15, No. 2/OPTICS EXPRESS 408).
A-scans are commonly recorded with 400 Hz to 400 kHz, in exceptional cases even in the MHz range. Ophthalmologic OCT systems have typical sensitivities of 80 dB to 110 dB. The utilized wavelength depends on the desired scan region and the absorption and scattering behavior of the tissue. Retinal OCT's frequently operate in the range of 700 nm to 1100 nm, while anterior chamber OCT's preferably use longer wave radiation, e.g., 1300 nm, which is absorbed in the vitreous humor. Anterior chamber OCT's can also be realized through switchover from retinal OCT's (US 2007/0291277 A1).
The axial measurement resolution of the OCT method is determined by the so-called coherence length of the applied light source, which is inversely proportional to the bandwidth of the applied radiation and typically lies between 3 μm and 30 μm (short coherence interferometry). The lateral measurement resolution is determined by the profile of the measuring beam in the scan region and lies between 5 μm and 100 μm, preferably below 25 μm. Due to its particular suitability for the examination of optically transparent media, this method is widespread in ophthalmology.
Two different basic types of OCT procedures have prevailed among those used in the field of ophthalmology. With the first type, the reference arm is modified in length to determine the measured data and continually measure the intensity of the interference without taking the spectrum into account. This procedure is called “Time Domain” procedure (U.S. Pat. No. 5,321,501 A). With the other procedure, called “Frequency Domain” procedure, however, the spectrum is taken into account for determining the measurements and the interference of the individual spectral components is recorded. Therefore, one is a signal within the time domain, and the other is a signal within the frequency domain.
The advantage of the frequency domain is the simple and quick simultaneous measuring, wherein complete information about the depth can be determined without requiring movable parts. This increases both stability and speed (U.S. Pat. No. 7,330,270 B2).
In frequency domain OCT, a further differentiation is made by obtaining the spectral information either with the use of a spectrometer (“spectral domain OCT”, or SD-OCT) or spectral tuning of the light source (“swept source OCT”, or SS-OCT).
The big technological advantage of the OCT is the decoupling of the depth resolution from the transverse resolution. In contrast to microscopy, this allows for the recording of the three-dimensional structure of the object to be examined. The purely reflective and thus contact-free measuring makes it possible to generate microscopic images of live tissue (in vivo).
DE 196 24 167 A1 describes a method for coherence biometry and tomography with increased transverse resolution. The position of light emitting points along a measurement path on the surface and on the inside of objects is measured using a measuring light beam of a short coherence interferometer. Short coherence interferometry in this context basically means that light with a short coherence length is used and the length to be measured in the measuring beam is determined such that the length in the reference beam is continuously changed until interferences occur, which is only the case when the two beam paths within the coherence length of the applied light are at equal length. As a result, the known length of the reference beam equals the sought length in the measuring beam.
With the method for coherence biometry, the entire depth of the measuring object in z-direction is detected with a measuring beam, while a whole series of such interferometric distance measurements (e.g., in x-direction) are carried out at adjacent points and assembled to an image with coherence tomography.
The light beam illuminating the object is shifted relative to the object, e.g., in x-direction after every A-scan, and so the object structure is scanned line by line in z-direction. These lines are subsequently assembled to a cross-sectional image (tomogram).
With the described solution, an equally good and high transverse resolution along the entire interferometric measurement path is achieved such that a suitable optical imaging of the (dynamic) focus in the object generated by the moved optical element ensures the simultaneous alignment of the optical lengths of reference beam path and measurement beam path right up to the (thus coherent) measuring focus.
Additional polarization-effective, optical components can be used for reducing reflection losses on the surfaces of the components and for optimizing the beam splitter of the described interferometer system.
An efficient optical coherence tomography system for quick, three-dimensional imaging is described in U.S. Pat. No. 7,145,661 B2. Light which is polarized using a polarization beam splitter is impinged into the OCT system, and so the OCT detector operates in a low-noise system.
When an eye is scanned, the system detector can simultaneously generate a pixel image with a low frequency component and a pixel image with a high frequency component of every point. While the pixel image with a low frequency component resembles that of an image realized with a scanning laser ophthalmoscope (SLO), the image with a high frequency component corresponds to that of a two-dimensional OCT image. Due to the pixel-to-pixel correspondence between the simultaneously recorded SLO and OCT images, the OCT image can be converted for the entire region pixel by pixel “on the fly” into a 3D image in accordance with the SLO image.
The described solution provides a system for realizing precise three-dimensional OCT images of the eye tissue in an extremely fast manner.
The not yet published document DE 10 2009 041 996.9 relates to an ophthalmological biometry or imaging system and a method for detecting and evaluating measured data, determining variables, distances and/or geometric relationships of eye structures. For optimizing the detection of the measured values, the measurement arrangement has, among others, a control circuit, which is formed by a control unit, an optical scan unit and a position sensor. In a disclosed embodiment, polarization is adjusted between measurement arm and reference arm of the OCT interferometer in order to ensure sufficient signal strengths preferably in all regions of the scan. The polarization adjustment can be made, for example, through rotated birefringent wave plates, motor-driven fiber paddles, birefringent or polarization-rotating liquid crystal modulators, or fast electro-optical polarization modulators.
US 2007/291277 A1 describes a further optical coherence tomography system, which is preferably based on a Mach-Zehnder interferometer. Once again, polarization is adjusted between measurement arm and reference arm of the interferometer in order to ensure sufficient signal strengths in all regions of the scan, wherein fiber paddles are used for said purpose. However, it is frequently not possible to detect all central and peripheral cornea regions evenly well with the adjustment of said fiber paddles. There is also the particular problem that depth-dependent changes in the polarization state of the backscattered light can also be observed in the cornea.
As is known, the strength of OCT signals depends on the adjustment of the polarization states of the superimposed light from sample arm and reference arm of the OCT interferometer. At equal polarization states, maximal interferences are achieved but with polarization states that are orthogonally oriented to one another (e.g., linearly or circularly), no interference signals can be detected.
Deviations between the polarization states can be caused, e.g., by different beam guidances in sample arm and reference arm, e.g., by mirrors in a periscope arrangement or polarization-effective, optical components. Birefringent samples, such as corneas, crystalline lenses, or retinal nerve fiber layers of the human eye can also cause said deviations.
According to the solutions from the known prior air, the polarization in OCT systems is in part or predominantly adjusted with adjustable optical elements which influence the polarization effect, wherein fiber paddles are used most frequently which are fiber loops that can be rotated manually or with a motor. With birefringence, fiber paddles have a similar effect as known wave plates.
However, despite the use of optical elements for polarization adjustment, the solutions known from the prior art are not capable of obtaining optimized signals from samples with depth-dependent, polarization-changing effect.
Another example is the measurements at the birefringent, retinal nerve fiber layer (RNFL), the thickness evaluation of which is very important for diagnosing glaucoma and glaucoma progression analysis. Elaborate polarization-resolved measurements using scanning laser ophthalmoscopes (SLO) are known. OCT signals can be recorded polarization-resolved using so-called polarization-sensitive OCT systems (PS-OCT), wherein separate detection paths are realized for orthogonally polarized light portions. Yasuno et al. (Optics Express Vol. 17, Iss. 5, pp. 3980-3996 (2009)) describes a PS-OCT system for the anterior chamber region which is based on the swept-source frequency-domain optical coherence tomography.
For example, the resulting interference spectra of two orthogonally polarized reference light portions which are each superimposed with the backscattered sample light are recorded separately in an SD-OCT system using two spectrometers.
Separate detection channels can also be realized by temporally fast multiplexing, using very fast polarization modulators, such as acousto-optical modulators (AOM), electro-optical modulators (EOM) or Piezo fiber stretcher.
It is disadvantageous that all these systems are very elaborate with regard to their beam guidance, polarization separation, separated detection or realization of high-frequency and high-precision modulation signals and do not achieve an optimal signal-to-noise ratio under certain circumstances.